Human (and other) autonomous nervous systems generate and conduct electrical signals to and from the heart muscle. These cardiac electrical signals can be monitored by an electrocardiogram (ECG) apparatus configured to receive electrical signals associated with the physical cardiac activity, which is manifested through electrical activity available generally on the surface regions of the thorax. In normal operation, an ECG can measure voltages associated with the nerve and muscles involved in cardiac activity. Thus, surface-adhering electrodes may receive electrical signals generated by cardiac activity. The electrodes may be positioned at various, specific, pre-determined locations on the body determined, for example, with reference to various electrical (e.g., resistive) models of the body or by empirical, clinical studies. In one example, the thorax can be modeled using parallel columns with two electrically conducting tissue paths. One path may include a relatively lower resistive blood path while a second path may include a relatively higher resistive tissue path. These paths may be employed to form circuits between ECG electrodes.
ECG signals and other biomedical signals may be measured as analog differential signals. Acquiring biomedical signals may involve selectively amplifying very small analog signals found in the same environment as larger common-mode signals. Both small and large signals may be present in an environment from which biomedical signals are acquired. The biomedical signals may be effectively found as additive values with respect to the measurement system. An intended measured signal may be enhanced by using differential measurement techniques since the unintended signals from the environment are normally found as common-mode signals. Thus, when subtracted using a differential amplifier, the unintended signals may be rejected or attenuated with respect to the desired signal. In practice, for ECG signals of interest, small differential analog signals may have a small amplitude (e.g., on the order of ±5 mV). Therefore, the signals typically require amplification by amplifiers capable of significant common mode rejection.
ECG electrodes employed in clinical use typically require skin preparation to achieve low impedance electrode coupling. Low impedance electrode coupling facilitates minimizing the variation of DC coupled ECG input amplifiers. However, conventional electrode impedance variations have been commonplace with DC resistance values in the 2 kΩ to 10 kΩ range. At higher frequencies, impedance may drop to a few hundred ohms, but signals of interest to an ECG are typically low frequency (e.g., less than 500 Hz).
Patient impedance measuring may be affected by impedance associated with the interface between an electrode and the patient. This impedance is often referred to as an electrode/skin impedance. The electrode/skin impedance, which may also be referred to as contact resistance, may be as high as 1 MΩ. Due to the high input-impedance of amplifiers associated with receiving electrocardiac signals, small differences (e.g., 10 KΩ) in skin/electrode impedance between electrodes can yield differential-mode signal amplitudes exceeding ECG signal amplitudes, which may negatively affect ECG analysis and diagnosis, by compromising accuracy of the measured values.